This invention generally relates to methods and apparatus for continuous, non-invasive monitoring of blood pressure.
A method for the determination of non-invasive beat-by-beat (continuous) systolic and diastolic blood pressure has long been desired in physiological monitoring. With this information, rapid changes in the physiological state of a patient can be much better managed. Automatic blood pressure cuffs can be used for this application, by inflating them as rapidly as is possible; however, this provides blood pressure data only at 1 to 2 minute intervals, and each inflation can be painful for elderly or hypertensive patients.
A very reliable technique for continuously measuring blood pressure is to insert a saline-filled catheter through the patient's vascular system to the point at which it is desired to perform the measurements. The catheter is connected to a pressure sensor, which measures the pressure in the vessel. An alternative method uses a catheter with a pressure sensor at the tip that directly senses the blood pressure. However, these techniques involve making an incision through the patient's skin and inserting the catheter into a blood vessel. As a consequence, this invasive procedure entails some risk of complications for the patient.
An indirect, non-invasive process for continuously measuring blood pressure is based on the pulse transit time (PTT) which is the time required for a blood pressure pulse from the heart beat to propagate between two points in the vascular system. One apparatus for this technique includes an electrocardiograph that senses electrical signals in the heart to provide an indication when a blood pulse enters the aorta. A pulse oximeter is placed on an index finger of the patient to detect when the blood pressure pulse reaches that location. The pulse transit time between the heart and the index finger is measured and calibrated to the existing blood pressure that is measured by another means, such as by the automated oscillometric method. Thereafter changes in the pulse transit time correspond to changes in the blood pressure. Generally, the faster the transit time the higher the blood pressure. Thus, changes in the pulse transit time can be equated to changes in the blood pressure.
However, the electrocardiograph (ECG) senses electrical signals in the heart, which do not indicate the point in time when the blood pressure pulse actually leaves the heart upon the mechanical opening of the aortic valve. A time interval of varying length, known as the cardiac pre-ejection period (PEP), exists between peaks of the QRS wave of the electrocardiogram signal and the aortic valve opening. The inability of prior pulse transit time-based monitors to account for the cardiac pre-injection period resulted in an inaccurate measurement of the pulse transit time and thus blood pressure.
In addition, changes in the compliance of the blood vessels also affect the pulse transit time. Chronic changes in arterial compliance occur due to aging, arteriosclerosis or hypertension. Arterial compliance can also change acutely due to neural, humoral, myogenic or other influences. Previous monitoring systems have been unable to separate changes due to compliance from changes due to blood pressure. As a consequence, some degree of inaccuracy has existed in calculating blood pressure from the variation of the pulse transit time.
Some previous investigators in the area of continuous, non-invasive blood pressure (CNIBP) monitoring have favored tonometric methods. These methods do not require a blood pressure cuff, but do require some small mechanical device that applies pressure to an artery, along with some kind of vibration sensor for tonometric pressure estimation. Such a device is described in published U.S. patent application Publ. No. 2003/0149369 A1. Devices of this type have proven unreliable in practice.
Still other previous investigators have used formulas derived from the Bramwell-Hill equation. These formulas relate blood pressure to measured arterial pulse wave velocity (PWV) and measured arterial blood volume. When the heart beats, it sends a pulse of pressure through the arterial system. This pulse propagates through the system by distending the elastic walls of the arteries, and this mechanism can be approximately represented by a linear wave equation. The basic expression for phase velocity of arterial pulse pressure waves is the same as that for the phase velocity of electrical waves propagating in a cable or transmission line, which has the same form of wave equation. The velocity is
                              v          p                =                              1            LC                                              (        1        )            where vp is the phase velocity of the arterial pulse pressure wave. For cable waves, C is capacitance and L is inductance. For arterial pulse wave propagation, C is called the compliance of the artery, defined as the derivative of the lumen area with respect to the pressure, and is given by
                    C        =                                            ⅆ              A                                      ⅆ              P                                =                                    2              ⁢              π              ⁢                                                          ⁢                              r                3                                      Eh                                              (        2        )            where A is the area of the arterial lumen, P is the pressure, r is the radius of the arterial lumen, h is the thickness of the arterial wall, and E is the modulus of elasticity of the arterial wall. This expression for C is derived in many standard textbooks, such as Hemodynamics, 2d edition, pages 96–98, by W. R. Milnor, published by Williams & Wilkins, 1989. L, which is the mass per unit length along the artery of the blood and represents the inertia of the blood that opposes the pressure pulse, is given by
                    L        =                  ρ                      π            ⁢                                                  ⁢                          r              2                                                          (        3        )            where ρ is the density of blood. If we substitute Eqs. (2) and (3) into Eq. (1), we get the famous Moens-Korteweg equation:
                              v          p                =                              Eh                          2              ⁢              r              ⁢                                                          ⁢              ρ                                                          (        4        )            If one substitutes Eq. (3) and the definition of C as a derivative into Eq. (1), we get what is called the Bramwell-Hill equation:
                              v          p                =                                                            π                ⁢                                                                  ⁢                                  r                  2                                            ρ                        ⁢                                          ⅆ                P                                            ⅆ                A                                                                        (        5        )            From this equation one can obtain a formula that linearly scales a measured arterial area into a pressure, with the slope based on measured PWV. If one puts all the pressure terms on one side and all the area terms on the other side (ρ is a constant) and then integrates, one gets:
                                          P            ⁡                          (              t              )                                -                      P            ⁡                          (              0              )                                      ≅                  ρ          ⁢                                          ⁢                      v            p            2                    ⁢                      ln            (                                          A                ⁡                                  (                  t                  )                                                            A                ⁡                                  (                  0                  )                                                      )                                              (        6        )            Equation (6) is an approximation because the pulse wave velocity is a function of the arterial radius, and therefore the lumen area, and in the derivation of Eq. (6) it was assumed that vp is a constant with respect to A. It is also possible to use a linearized version of Eq. (6), given by:
                                          P            ⁡                          (              t              )                                -                      P            ⁡                          (              0              )                                      ≅                                            ρ              ⁢                                                          ⁢                              v                p                2                                                    A              ⁡                              (                0                )                                              ⁢                      (                                          A                ⁡                                  (                  t                  )                                            -                              A                ⁡                                  (                  0                  )                                                      )                                              (        7        )            In order to use Eq. (6) or (7) to generate P(t), the arterial pulse wave velocity (PWV) vp and the arterial (i.e., lumen) area A must be measured. In addition, the initial values P(0) and A(0) must be obtained, and the blood density, ρ, must be replaced with a constant computed from cuff pressure data and ultrasound area and PWV measurements, all of which constitutes a calibration step.
In previous methods, PWV measurements were typically taken by observing the pulse transit time between two widely separated sites, such as the heart and a fingertip. The pulse arrival times at the measurement sites are typically determined by impedance plethysmography or pulse oximetry. In one known prior method (disclosed in U.S. Pat. No. 5,857,975), the time of the pressure pulse's origin at the heart is determined from an EKG signal. The required area measurements can also be obtained from plethysmography: volume measurements can be turned into areas by assuming a length. The initialization data can be obtained from a pressure cuff. The latter two measurements are not always used, however.
There are also methods that use only pulse wave velocity; because these methods do not measure blood volume/area, they must use empirical relationships between PWV and blood pressure rather than Eq. (6) or (7). One such method is described in published U.S. patent application Ser. No. 2003/0167012.
In the scheme described in U.S. Pat. No. 5,857,975, measurement of area and calibration are replaced by use of an assortment of seemingly arbitrary factors. (This only highlights the fact that schemes that estimate blood pressure on the basis of PWV alone have no physical justification.)
There are several grounds on which existing blood pressure calculations based on Eq. (6) might be criticized. One such criticism concerns the measurement of arterial blood volume. In particular, measurement of arterial area by plethysmography is confounded by the highly elastic nature of the veins. Impedance plethysmography measures total blood volume, and it is often applied to a limb such as the arm or leg. Since this measurement includes the venous blood volume, it cannot be reliably processed to produce arterial lumen area. If the venous blood volume did not change, then it might be possible to calibrate for this effect, but the venous blood volume is strongly affected by the subject's position, since hydrostatic pressures can cause pooling of blood in the veins, which are more highly distensible than the arteries.
Another criticism is that the PWV measurement resulting from measurement at widely separated sites is not directly applicable to Eq. (6). The PWV is being used in Eq. (6) to give information on the mechanical properties of the artery; however, the PWV depends not only on the elasticity of the arterial wall, but also on its thickness and the size of the lumen. Any PWV measurement that is made between widely separated sites is actually measuring PWV over a collection of branches of the arterial tree. Because these branches will have varying lumen area, the measured pulse transit time will reflect a composite of the component pulse wave velocities, and so it will not accurately reflect the mechanical properties of any particular arterial segment. It is possible to regard such a measurement as an approximation, but since Eq. (6) applies to a uniform, cylindrical tube, errors are inevitable, and their magnitude is hard to predict.
Another criticism of a method based on Eq. (6) concerns the way in which the PWV measurements are made. At every bifurcation and change in lumen radius along the arterial tree, a pulse wave reflection will occur. These pulse wave reflections can change the apparent (measurable) pulse transit time. The PWV that is needed in Eq. (6) is that which would be observed in a very long, uniform tube, where there would be no reflections.
A more fundamental criticism of blood pressure estimation based on Eq. (6) is that changes in physiological state can bring about changes in arterial wall elasticity. In general, this will change the measured PWV, and so the model of Eq. (6) will, to some extent, adapt to such changes. However, the operating point (P(0), A(0)) also depends on this elasticity, as does any multiplicative constant used to replace ρ. If the elastic modulus of the wall changes, then the same pressure will be associated with a different area. If the operating point and calibration constant are obtained using a pressure cuff calibration, then any changes of arterial elasticity would in principle require re-calibration, and if no such re-calibration is done, then gross errors in pressure estimation can occur. For example, if the mean pressure goes down while the elasticity goes up (which means that the elastic modulus E goes down), then the mean area could also go up. In that case, without an adjustment of (P(0), A(0)), the mean of the estimated pressure would go up rather than down.
There is a need for a method and means for acquiring data for use in the above-described blood pressure estimation scheme that overcomes the aforementioned drawbacks of prior art systems.